1. Field of the Invention
The present invention relates generally to nuclear medical (NM) imaging such as such as positron emission tomography (PET) and single photon emission computed tomography (SPECT), and more particularly to correction of emission contamination in transmission scans.
2. Description of the Background Art
Medical imaging is one of the most useful diagnostic tools available in modern medicine. Medical imaging allows medical personnel to non-intrusively look into a living body in order to detect and assess many types of injuries, diseases, conditions, etc. Medical imaging allows doctors and technicians to more easily and correctly make a diagnosis, decide on a treatment, prescribe medication, perform surgery or other treatments, etc.
There are medical imaging processes of many types and for many different purposes, situations, or uses. They commonly share the ability to create an image of a bodily region of a patient, and can do so non-invasively. Examples of some common medical imaging types are nuclear medical (NM) imaging such as positron emission tomography (PET) and single photon emission computed tomography (SPECT), electron-beam X-ray computed tomography (CT), magnetic resonance imaging (MRI), and ultrasound (US). Using these or other imaging types and associated machines, an image or series of images may be captured. Other devices may then be used to process the image in some fashion. Finally, a doctor or technician may read the image in order to provide a diagnosis.
PET uses positron-emitting isotopes to label common biological compounds in order to study metabolic and physiologic functioning. The biological compounds are injected into a patient and become concentrated in certain locations of interest in a patient's body. Shortly after a positron is emitted by the radioactive isotope, the positron collides with an electron, causing the positron and the electron to annihilate each other. Annihilation of the positron and electron results in a pair of 511 Mev gamma rays being emitted at the same time at approximately a 180° angle to each other. The patient is placed in a PET scanner to detect coincident emission of such gamma rays.
The PET scanner has an array of scintillation crystals and an array of photo-detectors for detecting when a gamma ray strikes a scintillation crystal. When two coincident gamma rays are detected, a record is made of the two scintillation crystals struck by the gamma rays. The two locations of the scintillation crystals define a line passing very close to the point of origin of the two gamma rays. A sufficient number of such coincident events are recorded in order to identify concentrations of the radioactive isotope in the patient. Digital image processing techniques permit the reconstruction of a three-dimensional image of the concentrations of the radioactive isotope in the patient.
For example, a commonly used positron-emitting isotope is Fluorine-18, which has a half-life about 110 minutes. The Fluorine-18 is produced in a cyclotron and is typically used to make Fluoro-deoxyglucose (FDG). FDG is a sugar that is metabolized by cells in the body. When FDG is injected into a patient, the FDG becomes distributed throughout the body in about an hour. The FDG, however, becomes more concentrated in parts of the body where the cells are more active. Since many cancers consist of very actively growing cells, the FDG becomes concentrated in such active cancers. When the patient is placed in the PET scanner, an image of the active cancers can be reconstructed from the coincident events recorded from the scanner. As described in Townsend et al. U.S. Pat. No. 6,490,476 issued Dec. 3, 2002, incorporated herein by reference, the PET scanner can be combined with an X-RAY CT scanner in order to provide anatomical images from the CT scanner that are accurately co-registered with the functional images from the PET scanner without the use of external markers or internal landmarks.
Image reconstruction from recorded coincident events can be more precise if corrections are made for scatter and attenuation of the gamma rays while the gamma rays pass through the patient's body. To enable such corrections, a transmission scan of the patient is made simultaneously with the recording of coincident gamma rays emitted from the patient. To make the transmission scan, the patient is irradiated by gamma rays from an external source, such as a Ge-68 (positron emitter) or Cs-137 (662 keV gamma) point source. However, any single gamma point source within the energy range of 50-700 keV can be used for the transmission scan. Recorded single gamma ray events presumed to originate from the point source are converted to an attenuation map of the patient. When a Cs-137 point source is used, the transmission data are scaled based on predetermined attenuation coefficients of the indicated class of body material; for example, for soft tissue or water, from a μ-value of 0.086 cm−1 at 662 keV to 0.095 cm−1 at 511 keV. The attenuation map is used to correct the associated emission scan of the patient.
Typically the recording of the transmission scan will cause emission contamination (EC) of the transmission data because of the presence of the 511 keV photons being emitted from the patient. The transmission data should be corrected for the emission contamination in order to avoid transmission image artifacts and underestimation of the reconstructed attenuation coefficients that in turn may result in incorrect attenuation and scatter coefficients. Several methods for performing such a correction are discussed in Hugo W. A. M. de Jong et al., “Correction for Emission Contamination in Transmission Scans for the High Resolution Research Tomograph,” IEEE Transactions on Nuclear Science, Vol. 51, No. 3, June 2004, pp. 673-676. These methods include histogram based scaling, segmentation, and subtraction of an estimated EC-contribution from the transmission data prior to reconstruction.
In histogram-based scaling the reconstructed grey-value of water is found by determining the water-peak in the mu-value histogram of the transmission image. Next, the image is scaled to the mu-value expected at 511 keV (0.095 cm−1). This method is said to intrinsically compensate for emission contamination. The transmission image can also be segmented to compensate for emission contamination. Segmentation is said to furthermore prevent the propagation of noise from the transmission image into the emission reconstruction. The threshold used for segmentation is set either to 50% of the theoretical expected mu-value of water (“fixed threshold”) or to 50% of the mu-value found by the histogram procedure.
It is also possible to subtract an estimate of the EC contribution from the transmission data prior to reconstruction. This EC-estimation is often measured using a so-called mock-scan, which essentially is a (fast) transmission scan without using the transmission source. De Jong et al. investigated two EC-estimation methods based on the assumption that the EC of a singles transmission scan has a relatively uniform spatial distribution. The EC was estimated by: i) a uniform flat distribution or ii) a general non-uniform distribution reflecting only the block-dependent sensitivity for EC. The non-uniform EC estimation was obtained by subtracting the transmission scans of a 6 cm cylindrical phantom filled with 37 MBq and without activity.